This invention relates to the field of cardiac prosthetic devices, and, more particularly, to hydraulically actuated total replacement artificial hearts and circulatory assist devices, including left ventricular assist devices, especially for use by and implantation in humans.
It has been estimated that between 16,000 and 50,000 patients annually are suitable candidates for implantation of a total cardiac prosthesis (TCP). Such candidates typically are disabled due to insufficient left and right ventricular function but are otherwise in good health. Many thousands more annually with inadequate left ventricular function and satisfactory right ventricular function may be candidates for a permanently implanted left ventricular assist device (LVAD).
The ideal total cardiac prosthesis must provide complete rehabilitation for the patient. Such a TCP recipient must be able to engage in gainful employment and all normal activities including moderate exercise. He should retain a substantially normal appearance and normal or near normal mobility with no significant limitations of any kind. Cardiac output effected by the TCP must be normal, adequate and sufficiently responsive to the patient's requirements to accommodate expected, sudden changes in physical activity or emotional stress level. The presence and operation of the TCP must be sufficiently unobtrusive so that the patient can largely forget that he is dependent on an artificial heart. All blood pumping functions of the TCP should be completely automatic, so that the patient performs no control or monitoring functions except for maintaining adequate power to the TCP, and responding to warnings that indicate a lack of power or serious problems requiring immediate technical or medical attention.
The intrathoracic blood pumping components of the TCP must be similar in size and weight to the natural heart. TCP life must be sufficiently long and reliability sufficiently high that risk to the patient of sudden prosthesis failure and its attendant anxiety are minimized. The formation of pannus and adherent thrombus must be prevented to avoid a compromise of blood pump function. Thrombo-emboli and excessive blood damage also must be prevented. The TCP must not damage adjacent tissues or impair organ function by toxicity and adverse tissue reactions, by mechanical trauma or compression, or by excessive local temperatures. The system must avoid skin penetrations of any kind to prevent infections that can arise from percutaneous leads. This eliminates a major risk to the patient, reduces the need for clinical observation and treatment, and reduces the maintenance of the TCP required of the patient. This ideal system must be low in cost to purchase, implant, and maintain. The frequency and extent of routine monitoring and maintenance, both medical and technical, must be low.
Serious research toward the realization of a total cardiac prosthesis has been under way since about 1957, sponsored largely by the U.S. National Institutes of Health (NIH). Researchers have directed this activity to six principal areas: (1) blood-compatible materials for the blood pumping means; (2) heart valves; (3) blood pumps; (4) blood pump actuating means; (5) power supplies and their application to the internal-blood pump actuating means; and (6) control mechanisms for the pumping function.
Many materials have been developed which apparently achieve blood compatibility. See, e.g., the recent survey and evaluation of these by M. Szycher et al in "Selection of Materials for Ventricular Assist Pump Development and Fabrication", Trans. ASAIO, Vol. XXIII, p. 116, 1977, incorporated herein by reference. (As used herein, Trans. ASAIO refers to the Transactions American Society For Artificial Internal Organs). While there is yet no human experience, recent materials like Biomer, Avcothane, etc., have been benign (i.e., have not caused thrombo-emboli) for periods up to 221 days in the calf. The materials for blood pumping membranes or sacs, however, must not only be benign and tissue compatible, but also able to withstand tens of pounds of force for something on the order of 10.sup.9 flexing cycles during a 20-year prosthesis life. Apparently, appropriate materials are nearly, if not already, realized today.
Another critical element of the blood pumping means is the valves, which permit blood flow into or out of the heart, but prevent backflow of blood. Many different types of valve prostheses have been developed and used in tens of thousands of implants to replace defective natural valves. Hence, adequate valves for a TCP appear to be well within the state of the art.
There has been a great deal of development activity in the area of blood pumps, primarily associated with LVAD's. This experience has shown that by utilizing appropriate biocompatible materials as described above, adequate and reliable blood pumps can be designed. The most common form of blood pump is comprised of an elastomeric sac or diaphragm-capped cavity. In a TCP which comprises two such blood pumps, each cavity is fitted with an inlet valve and an outlet valve. These pumping cavities replicate the function of the adjacent right and left ventricles of the natural heart.
The least developed of the aforementioned areas of activity is the development of an actuator to couple the power supply to the blood pump. In order to squeeze the blood-pumping sacs or force the diaphragm into the blood pumping cavity, pneumatic actuation means supplied from outside the body are most common. A number of mechanical actuation systems may be found in the literature. All sorts of linkages, gears, cams, etc., have been proposed, but none is known to be successful. Most of these systems are driven by an electric motor, although some have relied upon piezoelectric devices and other esoteric means. Both copulsation, the technique used by the natural heart, and alternate pulsation of left and right ventricles have been employed successfully. See Smith, L. M.; Olson, D. B., Sandquist, G., Grandall, E., Gentry, S., and Kolff, W. J., "A Totally Implantable Mechanical Heart", Proceedings from the European Society of Artificial Organs, Vol. 2, p. 150, 1975. Medical opinion appears to be impartial regarding this choice.
The coupling of a mechanical drive to the sensitive blood pumping diaphragm or sac is difficult to accomplish without raising excessive stresses and causing fatigue failures. Hence, the preferred coupling means is fluid, either liquid or gas. For example, one group has constructed an electric motor powdered, cam-actuated, diaphragm air pump which couples to the blood-pumping sac via pneumatic pressure. See V. Poirier et al, "Advances in Electrical Assist Devices," Trans. ASAIO, Vol. XXIII, p. 72, 1977, incorporated herein by reference. The entirety of the above-described mechanism is intended to be implanted within the thoracic cavity. In the Poirier et al design the motor rotates only once per heartbeat. Because relatively large torque is required from the motor, it must use strong magnetic fields, employ high current, and is rather heavy.
Burns et al, by contrast, constructed a TCP actuation system using a 10k-40k rpm motor driving a hydraulic pump pressurizing a liquid to actuate the blood-pumping bladders. See W. H. Burns et al, "The Totally Implantable Mechanical Heart, an Appraisal of Feasibility," Annals of Surgery, Sept. 1966, pp. 445-456, and W. H. Burns et al, "The Total Mechanical Cardiac Substitute," Process in Cardiovascular Diseases, Vol. XII, No. 3, 1969, pp. 302-311, both incorporated herein by reference. However, the electromechanically actuated hydraulic switching valve used in this and similar systems to shunt hydraulic fluid back and forth between ventricular actuating chambers has a number of disadvantages. The switching valve itself is relatively large and heavy, consumes a great deal of power and is potentially unreliable. Long and large ducts required in this type of system cause undesirable large frictional and inertial losses, and long fluid acceleration times.
Another approach to hydraulic actuation taken by researchers has involved the use of a reversible pump which directly pumps fluid back and forth between the two actuating chambers. See Jarvik U.S. Pat. No. 4,173,796.
On the subject of power, up to this time most TCPs implanted in the calf have been powered pneumatically via transcutaneous tubing into the thoracic cavity. A large external console supplies the proper regimen of pressure variations in order to activate the internal blood pump. With such a system, calves have lived up to 221 days. Jarvik, "The Total Artificial Heart", Scientific American, Vol. 244, No. 1, pp. 74-80, January, 1981. On another tack, NIH has sponsored considerable effort on the development of internal nuclear power supplies and, to a lesser extent, of chemical fuel cells. None of this work, however, appears to be promising; in fact, the nuclear effort was terminated by the U.S. Energy Research and Development Administration. Additionally, various means of transmitting mechanical power transcutaneously have been attempted, but none appears to be promising. At present transcutaneous transmission of electricity appears to be the preferred method for powering a TCP. A second, less preferable, possibility is the supplying of electrical power through percutaneous wire penetrations, but these always pose a threat of infection and are psychologically annoying to the patient.
Several investigators have developed the technique of transcutaneous electrical power transmission. Their approach is to implant a coil under the skin. This coil functions as a transformer secondary winding, receiving power from an inductively coupled, external, mating coil juxtaposed therewith to serve as the transformer primary winding. At frequencies on the order of 17 kHz, up to 100 watts have been thus transmitted for many months across the skin of a dog, by Schuder. See, J. C. Schuder et al, "Ultra High Power Electromagnetic Energy Transport Into the Body," Trans. ASAIO, 1971, incorporated herein by reference. Thus, it appears that the inductive delivery across the intact skin of the approximately 30 watts needed to power a TCP is well within the state of the art.
On the subject of control of a TCP to make it sympathetic to the body, there have been many different approaches and much controversy. Some researchers have attempted to provide no active control. Others have required a control in order to achieve regular beating. See, e.g., W. H. Burns et al, "The Total Mechanical Cardiac Substitute," identified above. Some systems have attempted to control systole (i.e., the contraction phase of the cardiac cycle whose rate is one determinant of cardiac output) from the left ventricle of the TCP in order to control the systolic pressure in the aorta. Still other systems have attempted feedback control of stroke volume and beat rate.
The natural heart and at least some, if not all, TCPs are comprised of two pumps in series. The right pump receives blood from the vena cava and impels it into the pulmonary artery. The left pump receives blood from the pulmonary vein and impels blood into the main circulatory system via the aorta. These two pumps must, over time periods considerably longer than that of a few beats, pump nearly the same amount of blood. Otherwise, the delicate pulmonary circuit will either collapse or rupture from a deficiency or excess of blood pumped by the right ventricle to the left. Various investigators have included controls in their TCP systems in order to achieve the critical balance between the pumping rate of the right and left ventricles. The major intrinsic mechanism by which the natural heart controls cardiac output is described by Starling's Law, which essentially states that a ventricle will expel during systole essentially that blood which flows into the relaxed ventricle during diastole. For the right ventricle, the body controls the "tone", i.e., the pressure in the venous system, so that the pressure in the vena cava (relative to atmospheric pressure) may rise from 5 to 15 mm Hg when there is a demand for higher blood flow. This pressure change causes approximately a proportional increase in the amount of blood which flows from the vena cava through the tricuspid valve into the relaxed right ventricle during diastole.
It is important to note that the natural heart has no means to suck upon the veins. It can only produce a systolic contraction which expels blood from the ventricular chamber.
Similarly, for the left ventricle, the pressure in the pulmonary vein varies from 5 to 15 mm Hg and produces a proportional increase in blood flow into the left ventricle. If the right ventricle should temporarily pump slightly more than the left ventricle, the pressure rises in the pulmonary artery, and, as a consequence, in the pulmonary vein, causing more blood to flow into the left ventricle and thereby matching the pumping rate of the left ventricle to that of the right ventricle. Thus, the natural heart achieves the necessary balance between the two pumps in series via simple and direct fluid dynamic means. In a real sense, the heart is the servant, not the master of the circulatory system, and in particular it responds in the final analysis to the requirements of the body as reflected by the peripheral oxygen saturation. The above-described intrinsic control can maintain body function even in the absence of extrinsic humoral or neural control.
The body also neurally controls the rate at which the natural heart beats. Cardiac output is a function of the amount of blood ejected during systole, and the rate at which the heart beats. For all but the most strenuous activity, the systolic stroke volume per beat remains substantially constant. Thus, cardiac output is primarily a function of beat rate (i.e., the number of beats per minute). Heart rates can vary from a low of about 40 to as high as 220 beats per minute in a young person and ordinarily from about 60 to 150 bpm in an adult. Cardiac output of the natural heart can vary from about 4 to as high as 24 liters per minute, the latter being the case of a trained athlete. Experience with pacemakers and transplanted natural hearts shows that beat rate control via neural sensors is unnecessary for a satisfactory life. The hundreds of natural hearts which have been transplanted operate at their own beat frequency, unresponsive to the body's neural demands because there is no neural connection.
The natural control system also ensures that the systolic pressure in the aorta does not drop below about 80 mm Hg, in order to maintain adequate circulation to the brain. The mean pressure in the aorta is established by cardiac output and the peripheral resistance of the vascular systems. In some of the TCPs which previously have been developed, a control means has been provided to maintain pressure in the aorta and atrium within a reasonable range. On the other hand, there is evidence from natural heart transplants that such control is unnecessary; transplanted human hearts have no neural connections to the host body and hence their systolic rates are not related to neural control, yet people with such transplants have been able to lead meaningful lives. It may be concluded that a TCP can be satisfactorily operated without such control. The evidence above teaches that a workable TCP can be made to approximate the natural heart's Starling's Law behavior with relatively simple control operations.
Thus, a TCP is now technically feasible provided that a competent design is constructed. The critical blood pumping technology appears to be well established and adequate for long-term survival of the recipient. Benign power transmission across the skin can obviate the portent of infection of the thoracic cavity transmitted via percutaneous leads. One major area where satisfactory progress is lacking, however, is the provision of a practical blood pump actuating mechanism. What is needed is a simple, lightweight, reliable, transcutaneously supplied, electrically-driven actuator. This objective is the one to which the present invention is principally addressed.